Implantable cardioverter defibrillator having a smaller mass

ABSTRACT

A capacitor-discharge implantable cardioverter defibrillator (ICD) has a relatively smaller mass of less than about 120 grams. The smaller mass of the ICD is achieved by selecting and arranging the internal components of the ICD to deliver a maximum defibrillation countershock optimized in terms of a minimum physiologically effective current (I pe ), rather than a minimum defibrillation threshold energy (DFT). As a result of the optimization in terms of a minimum effective current I pe , there is a significant decrease in the maximum electrical charge energy (E c ) that must be stored by the capacitor of the ICD to less than about 30 Joules, even though a higher safety margin is provided for by the device. Due to this decrease in the maximum E c , as well as corollary decreases in the effective capacitance value required for the capacitor and the net energy storage required of the battery, the overall displacement volume of the ICD is reduced to the point where subcutaneous implantation of the device in the pectoral region of human patients is practical. The size of the capacitor is reduced because the effective capacitance required can be less than about 125 μF. By optimizing both the charging time and the countershock duration for the smaller maximum E c , the size of the battery is reduced because the total energy storage capacity can be less than about 1.0 Amp-hours. In the preferred embodiment, the charging time for each defibrillation countershock is reduced to less than about 10 seconds and the pulse duration of the countershock is reduced to less than about 6 milliseconds.

This is a Continuation of application Ser. No. 08/412/920 filed Mar. 29,1995, now U.S. Pat. No. 5,827,326 which in turn is a continuation of08/263,257 filed Jun. 21, 1994, now issued as U.S. Pat. No. 5,405,363.

RELATED APPLICATIONS

This application is a continuation-in-part of the following co-pendingapplications, each of which are assigned to the assignee of the presentinvention, the disclosure of each of which is hereby incorporated byreference in this application:

U.S. patent application Ser. No. 07/910,611, filed Jul. 8, 1992,entitled "DEFIBRILLATOR PULSE GENERATOR", now issued as U.S. Pat. No.5,241,960;

U.S. patent application Ser. No. 07/835,836, filed Feb. 18, 1992,entitled "OPTIMAL PULSE DEFIBRILLATION METHOD AND IMPLANTABLE SYSTEMS";

U.S. patent application Ser. No. 07/953,485, filed Sep. 29, 1992,entitled "SHORT-PULSE IMPLANTABLE DEFIBRILLATION SYSTEMS";

U.S. patent application Ser. No. 07/808,722, filed Dec. 17, 1991,entitled "SMALL-CAPACITANCE DEFIBRILLATION PROCESS", now issued as U.S.Pat. No. 4,342,399;

U.S. patent application Ser. No. 07/808,722 filed Dec. 11, 1992,entitled "PROCESS AND APPARATUS FOR A DEFIBRILLATION SYSTEM WITH A SMALLCAPACITOR";

U.S. patent application Ser. No. 07/913,626, filed Jul. 16, 1992, as acontinuation of U.S. patent application Ser. No. 07/670,188, filed Mar.15, 1991, and entitled "DUAL BATTERY SYSTEM FOR IMPLANTABLEDEFIBRILLATORS", and issued as U.S. Pat. No. 5,235,979;

U.S. patent application Ser. No. 07/993,094, filed Dec. 18, 1992,entitled "STAGED ENERGY CONCENTRATION FOR A DEFIBRILLATOR", issued asU.S. Pat. No. 5,407,444; and

U.S. patent application Ser. No. 07/993,292, filed Dec. 18, 1992,entitled "SYSTEM AND METHOD FOR DELIVERING MULTIPLE CLOSELY SPACEDDEFIBRILLATION PULSES", now issued as U.S. Pat. No. 5,383,907.

TECHNICAL FIELD

The present invention relates generally to the field of automatic,implantable cardioverters and defibrillators. More particularly, thepresent invention relates to an implantable cardioverter defibrillator(ICD) that is a capacitor-discharge device having its internalcomponents, including a battery and a capacitor, selected and arrangedin such a manner that permits effective subcutaneous implantation of thedevice in the pectoral region of human patients.

BACKGROUND OF THE INVENTION

Existing implantable cardioverter defibrillators (ICDs) are typified bya relatively large size that usually requires implantation of theprosthetic device in the abdominal cavity of a human patient. In orderto allow for effective subcutaneous implantation of a prosthetic devicein the pectoral region of a human patient, the maximum size of theprosthetic device needs to be less than about 40-90 cc, depending uponthe physical size and weight of the patient. Unfortunately, all existingICDs have total displacement volumes of at least 110 cc or greater andenergy storage capacities of at least 2.0 Amphours. Even though thereare numerous advantages to developing an ICD having a displacementvolume small enough to permit implantation of the device in the pectoralregion of a human patient, to date it has been difficult to develop apractical ICD having a total displacement volume of less than about 100cc.

For reasons of simplicity and compactness, existing ICDs are universallycapacitor-discharge systems that generate high energycardioversion/defibrillation countershocks by using a low voltagebattery to charge a capacitor over a relatively long time period (i.e.,seconds) with the required energy for the defibrillation countershock.Once charged, the capacitor is then discharged for a relatively short,truncated time period (i.e., milliseconds) at a relatively highdischarge voltage to create the defibrillation countershock that isdelivered through implantable electrode leads to the heart muscle of thehuman patient.

One of the primary reasons why capacitor-discharge ICDs of a smallervolume have not been developed to date relates to the electricalrequirements for storing the high energy cardioversion/defibrillationcountershocks that are currently used to defibrillate human patients.Cardioversion countershocks have delivered energies of between about 0.5to 5.0 Joules and are used to correct detected arrhythmias, such astachycardia, before the onset of fibrillation. Defibrillationcountershocks, on the other hand, have delivered energies greater thanabout 3.0 Joules and are use to correct ventricular fibrillation or anadvanced arrhythmia condition that has not responded to cardioversiontherapy.

Presently, all capacitor-discharge ICDs are designed such that thecapacitor can store a maximum electrical charge energy of at least about35 Joules. In contrast, implantable pacemakers, which currently havedisplacement volumes of less than 50 cc, are designed to deliver pacingpulses of no more than about 50 μJoules. The requirement that acapacitor-discharge ICD be capable of storing an electrical charge withenough energy to deliver an electrical pulse almost one million times aslarge as that of an implantable pacemaker significantly increases thesize of the ICD over the size of the pacemaker due to the size of theelectrical components necessary to store this amount of electricalcharge energy.

The accepted requirement that ICDs be capable of storing a maximumelectrical charge energy of at least about 35 Joules arises out of thedefinition of an appropriate safety margin for the device according to aclinically developed defibrillation success curve as shown in FIG. 11.The defibrillation success curve plots the percentage probability ofsuccessful defibrillation for a ventricular fibrillation of about 5-10seconds versus the energy of a monophasic defibrillation countershock asmeasured in Joules. The safety margin for a given device for a givenpatient is presently accepted to be the difference between the maximumelectrical charge energy (E_(c)) stored by the capacitor in that deviceand the median defibrillation threshold energy (DFT) required for thatpatient.

Under existing medical practice, each time an ICD is implanted in ahuman patient, an intraoperative testing procedure is attempted in orderto determine the median DFT for that patient for the particularelectrode lead combination which has been implanted in the patient. Theintraoperative testing procedure involves inducing ventricularfibrillation in the heart and then immediately delivering adefibrillation countershock through the implanted electrode leads of aspecified initial threshold energy, for example, 20 Joules for amonophasic countershock. If defibrillation is successful, a recoveryperiod is provided for the patient and the procedure is usually repeateda small number of times using successively lower threshold energiesuntil the defibrillation countershock is not successful or the thresholdenergy is lower than about 10 Joules. If defibrillation is notsuccessful, subsequent countershocks of 35 Joules or more areimmediately delivered to resuscitate the patient. After a recoveryperiod, the procedure is repeated using a higher initial thresholdenergy, for example, 25 Joules. It is also possible that during therecovery period prior to attempting a higher initial threshold energy,the electrophysiologist may attempt to lower the DFT for that patient bymoving or changing the electrode leads.

The intraoperative testing procedure is designed to accomplish a numberof objectives, including patient screening and establishing a minimumDFT for that patient. Typically, if more than 30-35 Joules are requiredfor successful defibrillation with a monophasic countershock, thepatient is not considered to be a good candidate for an ICD andalternative treatments are used. Otherwise, the lowest energycountershock that results in successful defibrillation is considered tobe the median DFT for that patient. The use of the lowest energypossible for a defibrillation countershock is premised on the acceptedguideline that a countershock which can defibrillate at a lower energydecreases the likelihood of damage to the myocardial tissue of theheart. For a background on current intraoperative testing procedures,reference is made to M. Block, et al., "Intraoperative Testing forDefibrillator Implantation", Chpt. 3; and J. M. Almendral, et al.,"Intraoperative Testing for Defibrillator Implantation", Chpt. 4,Practical Aspects of Staged Therapy Defibrillators, edited byKappenberger, L. J. and Lindemans, F. W., Futura Publ. Inc., MountKisco, N.Y. (1992), pgs. 11-21.

Once the median DFT for a patient is established, theelectrophysiologist will determine a safety margin for a given ICDdevice usually by subtracting the median DFT from the maximum E_(c)stored by that device. Alternatively, a different calculation for thesafety margin is sometimes determined by estimating that point on thedefibrillation success curve where the electrical energy of adefibrillation countershock will insure a 99% success (E₉₉). Undereither definition, the safety margin needs to be large enough toaccommodate upward deviations along the defibrillation success curve.Such deviations may be expected, for example, with subsequent rescuedefibrillation countershocks delivered later in a treatment afterinitial cardioversion or defibrillation countershocks of lesser energieswere not successful. In these situations clinical data has found that,when delivered after 30 to 40 seconds of ventricular fibrillation, theelectrical energy necessary to achieve effective defibrillation mayincrease 50% or more over the median DFT. As a result, anelectrophysiologist usually will require that a given ICD have a firsttype of safety margin that is typically a factor of at least 2 to 2.5times the median DFT for that patient before the electrophysiologistwill consider implanting the given ICD in that patient. For thealternate E₉₉ point safety margin, the electrophysiologist will requirethat a given ICD have a maximum E_(c) at least 10 Joules above the E₉₉point.

Based on current clinical data that the average median DFT is somewherebetween 10-20 Joules for a monophasic countershock, the lower limit forthe maximum E_(c) that must be stored by the ICD is accepted to be atleast about 35 Joules, and more typically about 39 Joules, in order togenerate a maximum defibrillation countershock having an adequate safetymargin. The accepted lower limit for the maximum E_(c) of at least 35Joules is supported by clinical evaluations, such as Echt, D. S., etal., "Clinical Experience, Complications, and Survival in 70 Patientswith the Automatic Implantable Cardioverter/Defibrillator", Circulation,Vol. 71, No. 2:289-296, February 1985. In this article, the authorsevaluated data for early AICD devices having maximum E_(c) energies of32 Joules stored in a 120 μF capacitor with a discharge voltage V_(d) of750 Volts. In analyzing the clinical data for minimum DFTs, the authorsconcluded that the 32 Joule device had insufficient energy for effectivedefibrillation. It should be noted that in the next generation of theparticular AICD devices studied, the maximum E_(c) for the device (theCPI Ventak®) was increased to 39.4 Joules by increasing the capacitancevalue of the ICD by using a 140 μF capacitor.

Unfortunately, the requirement that an ICD be capable of storing amaximum E_(c) of this magnitude effectively dictates that the size ofthe ICD be greater than about 100 cc. This relationship between themaximum E_(c) that is required for an ICD and the overall size of theICD can be understood by examining how an ICD stores the electricalenergy necessary to deliver a maximum defibrillation countershock.

The only two components that impact on the ability of acapacitor-discharge ICD to store a maximum E_(c) are the capacitor andthe battery, which together occupy more than 60% of the totaldisplacement volume of existing ICDs. Thus, it will be apparent that thesize of a capacitor-discharge ICD is primarily a function of the size ofthe capacitor and the size of the battery. For a capacitor, the physicalsize of that capacitor is principally determined by its capacitance andvoltage ratings. The higher the capacitance value, the larger thecapacitor. Similarly, the physical size of a battery is also principallydetermined by its total energy storage, as expressed in terms ofAmp-hours, for example. Again, the higher the Amp-hours, the larger thebattery. With these concepts in mind, it is possible to evaluate how amaximum E_(c) affects the size of the capacitor and the size of thebattery in an ICD.

The maximum electrical charge energy (E_(c)) of an ICD is usuallydefined in terms of the capacitance value (C) of the capacitor thatstores the charge and the discharge voltage (V_(d)) at which theelectrical charge is delivered as defined by the equation:

    E.sub.c =0.5*C*V.sub.d.sup.2                               (Eq. 1)

The maximum electrical charge energy (E_(c)) can also be defined interms of how the energy is transferred from the battery to thecapacitor. In this case,, E_(c) is determined by the charging efficiency(e_(c)) of the circuitry charging the capacitor, the battery voltage(V_(b)), the battery current (I_(b)) and the charging time (t_(c)) asdefined by the equation:

    E.sub.c =e.sub.c *V.sub.b *I.sub.b *t.sub.c                (Eq. 2)

When Eqs. 1 and 2 are used to calculate a maximum E_(c) to be stored bythe device, the capacitance value (C) and the charging time (t_(c)) endup being the only true variables in these equations because theremaining values are all effectively determined by other constraints. InEq. 1, for example, the discharge voltage (V_(d)) for present ICDs canbe no more than about 800 Volts due to voltage breakdown limitations ofhigh power microelectronic switching components. As a result, V_(d) istypically between 650-750 Volts. In Eq. 2, it will be found that, forbatteries suitable for use in an ICD, the maximum battery output voltage(V_(b)) for ICDs is typically less than 6 Volts and, due to internalimpedances within these batteries, the maximum battery current (I_(b))is about 1 Amp. In addition, the charging efficiencies (e_(c)) ofexisting ICDs are presently on the order of about 50%.

When Eqs. 1 and 2 are evaluated for any given maximum E_(c), it will befound that there necessarily is a minimum capacitance value (C_(min))for the capacitor and a minimum charging time (t_(min)) required tostore that maximum E_(c) in the capacitor of the ICD. Knowing E_(c) andV_(d), Eq. 1 can be reworked as follows to solve for C_(min) : ##EQU1##

Similarly, knowing E_(c), V_(b), I_(b), and e, Eq. 2 can be reworked asfollows to solve for t_(min) : ##EQU2##

In other words, the fact that all ICDs presently use a maximum E_(c) ofat least 35 Joules means that all existing ICDs will require capacitorsof greater than 124 μF, and that all existing ICDs which draw 1 Amp ofcurrent from the battery will have a charging time of greater than 12seconds. Because the physical size of the capacitor is directlyproportional to the capacitance rating of the capacitor in farads for afixed voltage, the requirement that the capacitor be at least 124 μF iseffectively a minimum size limitation on the capacitor for dischargevoltages of less than about 800 Volts. Similarly, the requirement thateach charging time for a defibrillation countershock draw at least 12Amp-seconds of current from the battery is also a constructive minimumsize limitation on the battery. Thus, it can be seen that the existingrequirement for a maximum E_(c) of at least about 35 Joules effectivelydictates the size of both the capacitor and the battery and,consequently, the size of the ICD.

While existing ICDs have been successful in defibrillating humanpatients, and thereby saving lives, these devices are primarily limitedto implantation in the abdominal cavity due to their relatively largesize of greater than 110 cc. It has long been recognized that it wouldbe advantageous to reduce the total displacement volume of an ICDsufficiently to allow for subcutaneous implantation of the device in thepectoral region of human patients. This can only be done, however, solong as the device provides for a sufficient safety margin to insure itseffectiveness. Accordingly, it would be desirable to provide for anarrangement and configuration of the internal components of acapacitor-discharge ICD such that the total displacement volume of theICD is reduced, while a sufficient safety margin for the device isretained.

SUMMARY OF THE INVENTION

The present invention is a capacitor-discharge implantable cardioverterdefibrillator (ICD) having a relatively smaller displacement volume ofless than about 90 cc that permits effective subcutaneous implantationof the device in the pectoral region of human patients. The smallervolume of the ICD of the present invention is achieved by selecting andarranging the internal components of the capacitor-discharge ICD in sucha manner that the ICD delivers a maximum defibrillation countershockoptimized in terms of a minimum physiologically effective current(I_(pe)), rather than a minimum defibrillation threshold energy (DFT).One of the important results of optimizing the maximum defibrillationcountershock in terms of a minimum effective current I_(pe) is thatthere is a significant decrease in the maximum electrical charge energy(E_(c)) that must be stored by the capacitor of the ICD to less thanabout 30 Joules, even though a higher safety margin is provided for bythe ICD. Due to this decrease in the maximum E_(c), as well as corollarydecreases in the effective capacitance value required for the capacitorand the net energy storage required of the battery, the overalldisplacement volume of the ICD of the present invention is reduced tothe point where subcutaneous implantation of the device in the pectoralregion of human patients is practical.

By using a physiologically effective current (I_(pe)) to determine whatis a safe and effective maximum defibrillation countershock, the presentinvention takes advantages of the realization that it is the effectivecurrent delivered to the heart by the defibrillation countershock, andnot the total energy of the defibrillation countershock, that results ineffective defibrillation. In other words, the present inventionrecognizes that all Joules are not created equal and that the cells inthe heart muscle will make more effective use of some types ofelectrical energy and less effective use of other types of electricalenergy. The prior art technique of using a minimum DFT energy of thedefibrillation countershock to establish safety margins effectivelyignores the accepted fact that defibrillation countershock waveformswhich differ in shape, tilt and duration, for example, can havesignificantly different defibrillation threshold energies. In contrast,the effective current I_(pe) as used by the present inventionautomatically compensates for any differences in the effectiveness ofdifferent waveforms. Consequently, the ICD of the present invention usesa minimum effective current I_(pe) delivered to the heart muscle, ratherthan using a minimum DFT energy, as the measure for insuring an adequatesafety margin for the device.

To understand how the present invention can use a minimum effectivecurrent I_(pe) to insure an appropriate safety margin for the ICD, it isnecessary to recognize that the objective of any defibrillationcountershock is to generate an electric field across a large portion orall of the heart muscle, the myocardium. This electric field must have acurrent strong enough to extinguish all cardiac depolarizationwavefronts in the myocardium, and the current must be strong enough toprevent the myocardium cells from being restimulated during theirvulnerable period. In essence, the present invention recognizes that theelectric current generated by the defibrillation countershock must belarger than whatever minimum electric current is required for cellstimulation by at least a sufficiency ratio that will insure successfuldefibrillation. In this way, the use of an effective current I_(pe) canbe thought of as a correction factor applied to the actual current ofthe defibrillation countershock in order to compensate for the cellularphenomenon that currents below some minimum value simply do not have anyeffect on the cells.

It has long been known that in order to stimulate cells, a currentapplied to those cells must have a value at least equal to a rheobasevalue of those cells, otherwise the current applied to the cells is noteffective in stimulating the cells. C. Wiess, "Sur la Possibilite deRendre Comparable entre Eux les Appareils Suivant a l'ExcitationElectrique", Arch. Ital. deBiol., Vol. 35, p. 41 (1901); and L.Lapicque, "Definition Experimetelle de l'excitabilite", Proc. Soc.deBiol., Vol. 77, p. 280 (1909). Lapicque defined the rheobase value asthe stimulating current required for a pulse of infinite duration. Fromthis definition, he further defined a chronaxie value (d_(c)) to be theduration of a pulse that required a current twice that of the rheobasevalue. These two works have been combined in the literature to define astrength-duration model for the required average current for neuralstimulation known as the Weiss-Lapicque strength-duration curve, anexample of which is shown in FIG. 12.

The present invention builds on the Weiss-Lapicque strength-durationmodel to define a physiologically effective current I_(pe) as a simplemodel for the efficiency of a monophasic defibrillation countershock interms of the actual average current of the defibrillation countershock.The actual average current (I_(ave)) is given by the amount ofelectrical charge delivered at the electrode leads divided by theduration of the pulse delivering that charge. The end result of thederivation of a definition of effective current I_(pe) as taught by thepresent invention is that the effective current I_(pe) is given by thecharge delivered to the electrode leads divided by the sum of the pulseduration (d) and the chronaxie time constant for the heart (d_(c)).Expressing the charge delivered to the electrode leads in terms of theactual average current I_(ave) yields a definition equation as follows:

    I.sub.pe =(I.sub.ave *d)/(d+d.sub.c)                       (Eq. 5)

It can be seen from Eq. 5 that if the chronaxie value d_(c) were zero,the effective current I_(pe) would simply be I_(ave), the averagecurrent of a monophasic defibrillation countershock. In this way, thedefinition of an effective current I_(pe) distills the informationcontained in the Weiss-Lapicque strength duration curve to correct theactual average current I_(ave) of a monophasic defibrillationcountershock in order to compensate for the chronaxie phenomenon of thecells of the myocardium.

When a minimum effective current I_(pe) is used to select and arrangethe internal components of a capacitor-discharge ICD, the end result isa pair of surprising and non-intuitive conclusions.

First, the optimum capacitance value for the capacitor in acapacitor-discharge ICD is not determined by any stored or deliveredenergy requirement, but instead is a relatively constant value muchsmaller than any currently used capacitance values. The use of a minimumeffective current I_(pe) predicts that the optimum capacitance valuewill be a function of only the chronaxie time constant and theinter-electrode resistance of the electrode leads. This means that acapacitor with a smaller effective capacitance actually delivers adefibrillation countershock with more effective current I_(pe) than acapacitor having a larger effective capacitance. When the optimumcapacitance value is analyzed in terms of effective current I_(pe), itis found that the optimal capacitance value is given by the formula:

    C=(0.8*d.sub.c)/R                                          (Eq. 6)

Second, there is no single optimum pulse duration for a defibrillationcountershock having an arbitrary capacitance value. Instead, adefibrillation countershock of a shorter duration can provide a moreeffective current I_(pe) than a defibrillation countershock of a longerduration. The use of a minimum effective current I_(pe) predicts thatthe optimum pulse duration is a compromise between the RC time constantof the capacitor-discharge circuitry and the heart's defibrillationchronaxie time constant, d_(c). Thus, the predicted optimum pulseduration is not a constant, but rather is a function of the effectivecapacitance and other variables. The predicted optimum pulse durationcan be most simply, and robustly, expressed as a fixed tilt orexponential decay followed by a fixed time duration extension. When theoptimum pulse duration value is analyzed in terms of effective currentI_(pe), it is found that the optimal pulse duration is given by theformula:

    d=((R*C)+d.sub.c)/(e-1)                                    (Eq. 7)

Because the physical size of the capacitor is a function of itscapacitance rating, the use of a capacitor with a smaller effectivecapacitance provides for a significant reduction in the displacementvolume of the capacitor. In addition, because less energy is required tocharge up a capacitor with a smaller effective capacitance, a batterywith a smaller total energy storage, and, hence, a smaller displacementvolume, may also be used. Finally, the shortening of the duration of thedefibrillation countershock further decreases the energy requirements ofboth the capacitor and the battery, and also improves the safety marginof the device. In the preferred embodiment, several additionalinnovations are also used to further enhance the effectiveness of thedefibrillation countershock and decrease the energy storage requirementsof the ICD.

As a result of all of these improvements in the selection andarrangement of the internal components of the ICD of the presentinvention, the capacitor in the device only needs to store a maximumE_(c) of less than about 30 Joules, and preferably less than 27 Joules.The effective capacitance of the capacitor required by the presentinvention can be less than 120 μF, and preferably less than about 95 μF.By optimizing both the charging time and the countershock duration forthe smaller maximum E_(c), the size of the battery required by thepresent invention is reduced because the total energy storage capacityof the device can be less than about 1.0 Amp-hours. In the preferredembodiment, the charging time for each defibrillation countershock isreduced to less than about 10 seconds and the pulse duration of amonophasic defibrillation countershock, or of a first phase of amultiphasic defibrillation countershock, is reduced to less than about 6milliseconds.

By significantly reducing the displacement volume of both the capacitorand the battery, the overall displacement volume of an ICD in accordancewith the present invention can be reduced below 90 cc, and preferably tobetween 40-60 cc. Because the size requirements for effective pectoralimplantation will be distributed across the range from 40-90 cc for theentire population, it is obvious that the smaller the overalldisplacement of the ICD, the greater the percentage of human patientswho can benefit from pectoral implantation of the device. At thedisplacement volumes provided for by the present invention, subcutaneousimplantation of the device in the pectoral region of a human patient canbe quite practical and effective.

Accordingly, it is a primary objective of the present invention toprovide an implantable cardioverter defibrillator (ICD) having a smallerdisplacement volume than existing ICDs that permits effectivesubcutaneous implantation of the ICD in the pectoral region of humanpatients.

It is another primary objective of the present invention to provide anICD that delivers a maximum defibrillation countershock optimized interms of a minimum physiologically effective current (I_(pe)), ratherthan a minimum defibrillation threshold (DFT).

It is a further primary objective of the present invention to provide anICD with a discharge-capacitor that stores a maximum electrical chargeenergy (E_(c)) of less than about 30 Joules.

It is a still further primary objective of the present invention toprovide an ICD that utilizes a discharge-capacitor having an effectivecapacitance of less than 120 μF to store the electrical charge for thecardioversion/defibrillation countershock.

It is another objective of the present invention to provide an ICD thatdelivers a monophasic defibrillation countershock, or a first phase of amultiphasic defibrillation countershock, having a pulse duration of lessthan about 6 milliseconds.

It is a further objective of the present invention to provide an ICDthat has a battery and capacitor selected such that the battery cancharge the capacitor to its maximum E_(c) in less than about 10 seconds.

It is a still further objective of the present invention to provide anICD with a five year life and a battery having a total storage capacityof less than about 1.0 Amp-hours.

These and other objectives of the present invention will become apparentwith reference to the drawings, the detailed description of thepreferred embodiment and the appended claims.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is a frontal plan view showing the automatic, implantablecardioverter defibrillator of this invention implanted in the pectoralposition of a human patient.

FIGS. 2 and 3 are frontal and side plan views, respectively, of thepreferred embodiment of the ICD of the present invention.

FIGS. 4 and 5 are side and frontal plan views, respectively, showing thepower, capacitor, circuit and connector ports means positioned in thepreferred embodiment of the ICD of the present invention.

FIGS. 6 and 7 are plan views, showing the interior of the preferredembodiment of the ICD of the present invention.

FIG. 8 is a voltage versus time graph showing the relative distributionof the defibrillation energy discharged from the capacitor of thepreferred embodiment of the ICD of the present invention.

FIGS. 9 and 10 are voltage versus time graphs and a comparison table,respectively, showing the defibrillation energy discharged from thecapacitor used in the preferred embodiment of the ICD of the presentinvention versus the defibrillation energy discharged from a capacitorin a prior art ICD.

FIG. 11 is a defibrillation success curve used to define a minimumdefibrillation threshold (DFT) for prior art ICDs for monophasicintravenous defibrillation countershocks.

FIG. 12 is a typical Weiss-Lapicque strength-duration curve showing theaverage current required for defibrillation as a function of the pulseduration.

FIG. 13 is a defibrillation success curve for the present inventionusing an physiologically effective current (I_(pe)) for monophasicintravenous defibrillation countershocks.

FIG. 14 is a graph of the minimum effective current (I_(pe)) formonophasic intravenous defibrillation countershocks versus fibrillationtime showing the impact of prolonged fibrillation on minimum I_(pe).

FIG. 15 is a graph of the minimum effective current (I_(pe)) formonophasic intravenous defibrillation countershocks as a function ofelectrode resistance for both fixed duration and tilt countershockpulses.

FIG. 16 is a block diagram of a dual battery system energy storagesystem for the preferred embodiment of the present invention.

FIG. 17 is a block diagram of a rechargeable version of the dual batterysystem shown in FIG. 16.

FIG. 18 is a simplified circuit diagram of a prior art implantabledefibrillator circuit.

FIG. 19 is a simplified schematic circuit diagram of a staged energyconcentration circuit of the preferred embodiment of the presentinvention.

FIG. 20 is a simplified schematic circuit diagram of an alternateembodiment staged energy concentration circuit of FIG. 18.

FIG. 21 is a simplified schematic circuit diagram of another alternateembodiment staged energy concentration circuit of FIG. 18.

FIG. 22 is a schematic circuit diagram illustrating representative priorart circuitry for an implantable cardioverter defibrillator.

FIG. 23 is a schematic circuit diagram of one embodiment of theimplantable cardioverter defibrillator rapid pulse circuitry of thepreferred embodiment of the present invention.

FIG. 24 is a schematic circuit diagram of another embodiment of theimplantable cardioverter defibrillator rapid pulse circuitry of FIG. 23.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENT

In describing the present invention, first a description of thepreferred mechanical arrangement of the internal components of theimplantable cardioverter defibrillator (ICD) will be presented toprovide a context for the remainder of the description. Next, amathematical explanation of the derivation of the physiologicallyeffective current (I_(pe)) as used by the present invention will bepresented. Then, each of the major features responsible for decreasingthe overall displacement volume of the ICD will be described. Thesefeatures include: the use of a more optimal pulse duration and pulsewaveform for the cardioversion/defibrillation countershock, the use of acapacitor having a smaller effective capacitance value, and the use ofan improved battery configuration having a smaller total energy storage.

Mechanical Arrangement of the ICD Components

FIG. 1 shows an automatic ICD 17 of the present invention implanted inthe pectoral region 18 of the chest 11 of patient 10. The ICD 17 has aplurality of connector ports for connection to various implantablecatheter and other electrode means, as is known in the art. For example,electrode leads 41 and 42 are shown extending form the ICD 17 tocatheter electrodes 40 and 15 which are passed, respectively, into thesuperior vena cava 14 and the right ventricle 13 of heart 12. Further,lead 43 is shown extending from the ICD 17 to a subcutaneous patchelectrode 16. The specific configuration of the electrodes of thedefibrillation system is dependent upon the requirements of the patientas determined by the physician.

FIGS. 2 and 3 show the ICD 17 comprised of a housing 19 having matinghalf shells 21 and 22. Positioned and mounted on top of housing 19 is atop connector portion 20 having a plurality of connecting ports 23 whichare described further below. Importantly, the ICD 17 is comprised of acompact, self contained structure having predetermined dimensions whichpermit pectoral implantation. The housing 19 and top connector 20 areconstructed and arranged to yield a cooperating structure which housespower means, control means and capacitive means. This cooperatingstructure permits subcutaneous implantation in the pectoral region of ahuman patient and provides a compact and effective ICD thatautomatically senses the bioelectrical signals of the heart and is ableto provide a 750 volt capacitive discharge, for example, to the heartfor defibrillation purposes.

In the past, ICDs have required a size and configuration for functionalpurposes that necessitated implantation in the abdominal cavity of apatient. Such implantation has resulted in patient discomfort. However,the physical parameters of these prior art devices have preventedpectoral implantation, which is preferred by physicians and patientsalike. Table 1 below shows the size and weight comparisons between knownprior art ICD devices and the ICD 17 of the present invention.

                  TABLE 1    ______________________________________                       Present   Present Present             Prior Art Device    Device %                                         Device %             Device % of                       % of total                                 of Prior Art                                         of Prior Art             total     Device    Devices Devices             (by volume)                       (by volume)                                 (by volume)                                         (by weight)    ______________________________________    Connector              10        8        30      32    Capacitors              30        38       63      62    Batteries              30        23       38      57    Electronics              30        31       50      40    Total    100%      100%      50%     55%             (120 CC)  (60 CC)    ______________________________________

As shown in Table 1, the ICD 17 of this invention, provides a structurewhich is 50% of the volume of prior art devices and which has a weightwhich is 55% of the weight of the prior art devices. The connector port,capacitor, battery and electronic circuitry of the ICD 17 of the presentinvention are further described below.

It is important in this invention that the ICD 17 be constructed andarranged to minimize the overall displacement volume of the device toallow for pectoral implantation, for example. The housing structure 19is a compact and lightweight structure made of a biocompatable materialand has a contoured configuration. The overall structure of thisinvention has a weight of less than 130 grams, and preferably less than120 grams, and a volume of less than 90 cc, and preferably between about40-60 cc. As shown in Table 1, the ICD 17 of this invention hasgenerally 55% of the weight of prior art devices and a volume which isgenerally 50% of that of prior art devices. Table 1 further shows theweights and volumes of the respective components of this invention(connector, capacitor, batteries and electronics) as a percentage inweight and volume of the total and in comparison to prior art devices.

As further shown in FIGS. 2 and 3, the housing structure 19 has acontoured periphery which is matingly connected to the top connectormember 20 which also has a mating contoured configuration. The housing19 is constructed of a biocompatable material such as a titanium or astainless steel alloy. The top connector member 20 is also constructedof a biocompatable material, such as a biocompatable polymericcomposition. It has further been found that for pectoral implantationpurposes, that the housing structure 19 have a desired length to widthto thickness ratio of approximately 5 to 3 to 1.

When selected in accordance with the optimized minimum physiologicalcurrent (I_(pe)) as described below, the capacitor has an effectivecapacitance of approximately 85 uF, is constructed and arranged todeliver an initial discharge voltage V_(d) of 750 Volts, yielding theeffective defibrillation countershock which is also described below. Inthe preferred embodiment, the effective discharge voltage andcapacitance is achieved by using two flash-type capacitors in series,each having a capacitance rating of 170 μF and a voltage rating of 375Volts, while occupying a total displacement volume of only 7 cc each.The output of the capacitors is in communication with an electroniccircuitry output portion that generally is comprised of a flash typecircuit which delivers the capacitor discharge through electrodes 15, 16and 40, for example.

FIGS. 4 and 5 show the canister housing 19 having an interior space 30wherein capacitors 26 and 27 are positioned and wherein a battery system28 and circuit board portions 31 and 32 are positioned. The topconnector 20 is shown mounted to the top of the canister housing 19.Connecting ports 36, 37, 38 and 39 are shown positioned in the topconnector 20. The connector ports 36 and 37 are connectible to thepositive defibrillating electrode, for example, while connecting port 38is connectible to the negative defibrillating electrode, for example,and the connecting port 39 receives the pacing/sensing electrode leads41, 42. Channels 24 and 25 provide communicative and fastener membersthat provide for the attachment of the top connector 20 to the canisterhousing 19 and for the electrical connection between the ports 36, 37,38 and 39 and the electronic elements positioned in the interior space30 of housing 19.

As discussed, the top connector 20 of the defibrillator ICD 17 has, forexample, connecting ports 36 (DF+), 37 (DF+), 38 (DF-) and 39(sensing/pacing). The lead connected to the DF- port, for example, is inconductive contact with the catheter electrode 15 placed in the rightventricle 13 of the heart 12. The electrode lead(s) connected to the DF+port(s) are connected to either or both of the electrodes positioned inthe superior vena cava 14 and the subcutaneous patch electrode 16.Alternatively, the DF+ port holes may not be utilized, and plugged by astopper means, for example, when the ICD body itself is utilized as thepositive element to complete the defibrillation circuit. Thepacing/sensing electrode 44 provides an input to connecting port 39 ofthe ICD 17 and provides continual monitoring of cardiac signals from theheart. The circuitry of the ICD 17 has means to detect any tachycardiacor other arrhythmia condition and to thereby respond by the selectivedischarge of electrical energy stored in the capacitors 26 and 27.

As described in more detail below, the ICD 17 of this invention providesa device which utilizes smaller capacitors and batteries than those ofprior art devices and thus yields a countershock generator device havinga smaller displacement volume that permits effective implantation of thedevice in the pectoral region of a human patient. Although the smallerunit and associated components are smaller and deliver a smaller energycountershock to the heart, the implantation of the device in thepectoral region provides for a better countershock vector. Together withthe improved countershock pulse waveform as described below, the ICD 17produces a more effective defibrillation/ cardioversion countershockthan prior art ICD devices.

FIGS. 6 and 7 show the mating housing half shells 21 and 22,respectively of canister housing 19. The half shell 22 is shown to havean interior peripheral band 34 which is fixed adjacent the peripheraledge 33. The interior peripheral band 34 extends outwardly from the edge33 of half shell 22 and is constructed and arranged to receive theperipheral edge 35 of housing half shell 21. Alternatively, theperipheral band 34 may be mounted within housing half shell 21, wherebythe half shell 22 is positioned thereabout. The peripheral band 34 isalso provided to shield the electronic components within housing 19during the welding process uniting the body shells 21 and 22.

The flexible circuit board 29 is mounted within the interior space 30 ofhousing 19. The circuit board 29 provides for the sensing/pacingcircuitry in communication with the lead extending from connecting port39, for example. When a fibrillation episode is detected, the circuitboard 29 causes the capacitors 26, 27 to discharge an initial 750 Voltcharge through the electrode leads connected to ports 36-38, forexample, and to the heart 12 of the patient 10. The electronic circuitryhas a sensing portion which monitors the heart beat rate irregularity bymeans of two small electrodes 44, as is known in the art. In thepreferred embodiment, the circuitry further has a processor portionwhich determines, with respect to a predetermined standard, when theoutput portion of the circuit will be activated.

FIG. 8 is a graph showing the voltage discharge with respect to timefrom the 85 μF capacitor used in the preferred embodiment of the ICD 17of this invention. The graph shows the incremental benefit of thevoltage discharge with respect to time. FIG. 9 is a graph which showsthe instantaneous voltage with respect to time and compares the plottedvalues of a countershock having the same delivered energy content forboth the present invention and a typical prior art ICD. In FIG. 9, thecountershock is a 20 Joule delivered energy monophasic countershock andit will be seen that the pulse duration of the countershock inaccordance with the present invention is significantly shorter than thepulse duration of the countershock delivered by the prior art ICD.

As summarized in the table of FIG. 10, the 85 uF capacitor of thepreferred embodiment of the present invention provides 25.33 Joules ofdelivered energy in the form of a biphasic defibrillation countershockhaving a delivery efficiency of 97.5% from a 26 Joule maximum E_(c)stored in the capacitors 26, 27. In comparison, the 140 uF capacitorused in a prior art ICD device provides 34 Joules of delivered energy inthe form of a monophasic defibrillation having a delivery efficiency of86.3% from a maximum E_(c) stored in the capacitor of 39.4 Joules. Theeffective current I_(pe) of the biphasic countershock delivered by thepresent invention is 5.67 Amps, uncorrected, and possibly as high as6.55 to 7.32 Amps, when corrected to be a monophasic equivalent current.In contrast, the effective current I_(pe) of the monophasic countershockdelivered by the prior art device is 6.79 Amps. Thus, the uncorrectedI_(pe) of the present invention is only 20% less than the I_(pe) of theprior art device, while the maximum E_(c) of the present invention ismore than 50% less than the maximum E_(c) of the prior art device.

When the corrected I_(pe) provided by the biphasic countershock of thepreferred embodiment of the present invention is compared, the presentinvention provides essentially the same effective current I_(pe) as theprior art device with half the maximum E_(c) and, as little as half therequisite displacement volume for the capacitor. Depending upon thecorrection factor applied to convert the current efficiency of anoptimized biphasic countershock pulse to a traditional monophasiccountershock pulse (a 25% more energy efficient countershock is a 15%more efficient effective current, whereas a 40% more energy efficientcountershock is a 28% more efficient effective current), the correctedI_(pe) of the present invention is between 3% less to 7% more than theI_(pe) of the prior art device.

Derivation of the Physiologically Effective Current (I_(pe))

The famous Weiss-Lapicque model was developed at the turn of thecentury. It was an empirical model and the first physiologicalexplanation for why the model accurately predicts the required currentfor cellular stimulation was only recently explained. Irnich, W., "TheFundamental Law of Electrostimulation and its Application toDefibrillation", PACE 1990, 13 (Part 1): 1433-1447. The model gives therequired (average) current for neural stimulation as:

    I.sub.ave =K.sub.1 +(K.sub.2 /d)                           (Eq. 8)

with d being the pulse duration. The value K₁ is the current requiredfor an infinite duration pulse. The "chronaxie" is that duration whichrequires a doubling of the rheobase current. The chronaxie time constantd_(c) is thus given by:

    d.sub.c =K.sub.2 /K.sub.1                                  (Eq. 9)

Defining I_(r) as the rheobase current gives:

    I.sub.ave =I.sub.4 * 1+(d.sub.c /d)!                       (Eq. 10)

The Weiss-Lapicque model was based on cell stimulation, notdefibrillation. However, in 1978, Bourland et al showed, with a study ofdogs and ponies, that defibrillation thresholds also followed theWeiss-Lapicque model when current averaged over pulse duration was used.Bourland, J. D., Tacker, W. A. and Geddes, L. A., "Strength DurationCurves for Trapezoidal Waveforms of Various Tilts for TranschestDefibrillation in Animals", Med. Instr. (1978), Vol. 12, No. 1:38-41. Atypical strength-duration curve is shown in FIG. 12.

Bourland et al. further proposed that the average current of a pulse wasthe best measure of its effectiveness when compared to other pulses ofthe same duration. This was found to hold fairly true for pulses from2-20 ms in duration, regardless of waveform. Bourland, J. D., Tacker, W.A. and Geddes, L. A., et al. "Comparative Efficacy of Damped Sine Waveand Square Wave Current for Transchest Ventricular Defibrillation inAnimals", Med. Instr. (1978), Vol. 12, No. 1:42-45.

Numerous studies have confirmed the strength-duration relationship fordefibrillation currents. These same studies show that the defibrillationchronaxie time constant d_(c), is in the range of 2-4 ms. Using theavailable data on measured defibrillation chronaxie time constants,d_(c) =2.7±0.9 ms is the average chronaxie value for the human heart.

In contrast to the accepted prior art technique of using a minimumdefibrillation threshold energy (DFT) to measure the effectiveness of adefibrillation countershock, or even in contrast to the suggestion byBourland et al to use the average current, the present invention definesan effective current as that percentage of the rheobase requirement forthe human heart that the average current of a defibrillationcountershock pulse can satisfy. Under this definition, successfuldefibrillation will require that

    I.sub.ave >=I.sub.r * 1+(d.sub.c /d)!                      (Eq. 11)

where I_(ave) is the current averaged over the pulse duration of thedefibrillation countershock. Satisfying this condition and substitutingI_(pe) for I_(r) yields a definition of physiologically effectivecurrent (I_(pe)) which can be expressed in several ways: ##EQU3##

Note that the effective current of a defibrillation countershock onlyequals the rheobase current when the output of the pulse is exactlyoperated at the defibrillation threshold, and, hence, with a zero safetymargin. In general, the two parameters are not equal in value ororientation. The effective current I_(pe) is a system variable of theICD, while the rheobase current I_(r) is primarily a physiologicvariable.

FIG. 13 shows a defibrillation success curve for monophasic intravenousdefibrillation countershocks plotted in terms of the effective currentI_(pe) of the present invention. It will be apparent when comparing theI_(pe) defibrillation success curve shown in FIG. 13 with the DFTdefibrillation success curve shown in FIG. 11 that the I_(pe) curve istighter and the necessary safety margin is much closer to the medianI_(pe) required for effective defibrillation. FIG. 14 is a graph of theminimum I_(pe) for monophasic intravenous defibrillation countershocksversus fibrillation time showing the impact of prolonged fibrillation onminimum I_(pe). Together, these figures illustrate how a device with asmaller maximum E_(c) can still provide a more than adequate safetymargin, as long as the effective current I_(pe) of the defibrillationcountershock is sufficient. The next two sections of the description setforth how to optimize the characteristics of acardioversion/defibrillation countershock in terms of effective currentI_(pe).

Optimal Pulse Waveform and Duration

The present invention uses the effective current I_(pe) model to findthe optimum pulse duration for a conventional, time-truncated, capacitordischarge defibrillation countershock waveform. In such acapacitor-discharge system, a capacitance (C) is charged to an initialvoltage (V_(i)) and then discharged into the effective load resistance(R) of the heart for a pulse duration (d), at which time the capacitancewill have a final voltage (V_(f)). The amount of droop in thecapacitor-discharge waveform at the time the pulse is truncated has beenreferred to as the "tilt" of the waveform as given by the equation:

    tilt=(V.sub.i -V.sub.f)/V.sub.i                            (Eq. 15)

    =1-V.sub.f /V.sub.i

Substituting the RC time constant exponential decay for V_(f) /V_(i)yields:

    tilt=1-e.sup.-d/RC                                         (Eq. 16)

where RC is the capacitor-discharge system time constant, also known asτ. Substituting the system parameters for C, V and tilt for thedelivered charge in Eq. 14, the effective current I_(pe) can also beexpressed as:

    I.sub.pe =(C*V*tilt)/(d.sub.c +d)                          (Eq. 17)

    =(C*V*(1-e.sup.-d/τ))/(d.sub.c +d)

Because the first derivative of I_(pe) approaches zero at extreme valuesof d, its maximum is at the point of zero derivative. ##EQU4##

Normalizing Eq. 18 to the system time constant by defining:

    z=d/τ                                                  (Eq. 19)

    α=d.sub.c /τ                                     (Eq. 20)

The derivative of.Eq. 18 now reduces to

    0=(z+α+1)e.sup.z -1                                  (Eq. 21)

Multiplying by -e^(z) and defining f(z) gives:

    0=e.sup.z -z-α-1≡f(z)                          (Eq. 22)

This transcendental equation cannot be solved in closed form, so theNewton-Raphson approximation is used. If z₀ is the first approximationfor the root, then z'=z₀ -(f(z₀)/f'(z₀)) is the Newton-Raphsonapproximation for Eq. 22. Present ICD devices favor a tilt of about 65%.This implies that the countershock pulse duration is roughly equal toone system time constant (d≈τ). Because z=d/τ≈1, the first approximationis z₀ =1. Thus: ##EQU5##

Denormalizing Eq. 24 gives:

    d≈τ (1+(d.sub.c /τ)/(e-1)!                 (Eq. 25)

This gives the expression for optimal pulse duration of a time truncatedmonophasic capacitor discharge:

    d=(τ+d.sub.c)/(e-1)                                    (Eq. 26)

Numerical optimization shows that this estimate gives an I_(pe) within0.2% of true optimum for typical values of R, C and d_(c). For extremevalues the maximum error in the resulting I_(pe) is less than 2.0%. Itshould also be noted that (e-1)≈1.72 ≈2. Thus, the optimal pulseduration, d, is approximately equal to the average of the capacitor timeconstant and the heart's chronaxie time constant. In other words, themodel suggests that the best pulse duration is a compromise between thetime required to deliver the capacitor's charge, τ=RC, and the timerequired to match the timing of the heart, d_(c).

The predicted optimum d may be used to directly derive the optimum tiltfor the countershock pulse from Eqs. 16 and 26.

    tilt.sub.opt =1-e.sup.-d.sbsp.opt /τ                   (Eq. 27)

    =1-exp - (1+(d.sub.c /τ))/(e-1)!!

where, again, τ=RC. Assume, for a moment, that RC is chosen to equald_(c). From Eq. 27 we have that tilt_(opt) =68.8%.

It will be noted that the predicted optimum tilt is rather high forsmall capacitance values. An intuitive explanation is that they need tospread their charge delivery over as great a duration as possible tocome closer to the chronaxie time. Because the rheobase is smallcompared to the peak current, this lowering of average current is welltolerated. Conversely, for large capacitances the optimum tilt issmaller as the average current must remain above the rheobase.

The prediction by this model that small capacitor systems benefit fromhigher tilts is supported, but not predicted, by a prior study oftruncated waveforms. Schuder, J. C. et al. "Transthoracic VentricularDefibrillation in the Dog with Truncated and Untruncated ExponentialStimuli", IEEE Trans. Bio. Eng., (1971); BME-18:410-415. This studyshows that the greatest improvements from truncation are obtained withpulses with d>5 ms. In other words, high tilts are better tolerated withsmaller pulse durations (which are generated by smaller capacitors).

It will be understood that the inter-electrode resistance varies withthe patient and positioning of the electrode leads and may change afterimplantation of the ICD. It is desirable that the pulse duration dremain close to optimum in spite of this change. FIG. 15 gives theeffective current I_(pe) for various inter-electrode resistances whenthe tilt and duration of the pulse were optimized for an assumed 50 Ωload. In this example, a 140 μF capacitor and a 2.7 ms chronaxie areassumed. One curve shows how the effective current I_(pe) varies whentilt is used as the specification for the pulse duration, while theother reflects the use of a fixed time duration. Note that tilt besttolerates decreases in resistance while a fixed duration best handlesincreases in resistances.

The optimum pulse duration from Eq. 26 may be rewritten as:

    d=0.58 RC+0.57 d.sub.c                                     (Eq. 28)

Because 1-e⁻⁰.58 32 44%, a pulse duration of 0.58 RC may be redefined asa 44% tilt. Thus, the optimum duration from Eq. 28 may be stated inwords as:

1. Allow the capacitor voltage to decay by 44%, then

2. Continue the pulse for an additional 58% of the chronaxie timeconstant.

If we assume d_(c) =2.7 ms, then this gives a 44% tilt followed by a 1.6ms extension for optimum pulse duration. Note that this specification ofthe optimum pulse duration automatically adjusts the duration for anychanges in resistance so that continued monitoring of the effectiveinter-electrode resistance is not required. In addition, using thisspecification of the optimum pulse duration gives an I_(pe) in excess ofthat specified by either a fixed tilt or a fixed duration alone, for anyresistance value.

The choice of a tilt or duration specification has been an open issue indefibrillation. Based on the predictions using the effective currentmodel of the present invention, it would appear that the best choice isactually a composite of tilt and duration as given in Eq. 28. Thisspecification for pulse duration is intuitively attractive in that itdirectly recites the necessary compromise between the electronics(duration sufficient for charge delivery) and the heart (duration closeto chronaxie for efficiency).

The optimized pulse duration utilized by this invention can be appliedto monophasic waveforms, or the first phase of a biphasic or multiphasicwaveform. In the latter cases, when it is applied to the first phase ofa waveform, subsequent phases of the waveform can be specified to haveequal or lesser duration than the first phase. In the preferredembodiment, a biphasic waveform is used in order to take advantage ofthe additional decrease in the minimum effective current I_(pe) requiredfor effective defibrillation that is predicted by known clinical datashowing a 25-40% decrease in minimum DFT energy thresholds for biphasiccountershock pulses.

It will also be appreciated that the optimum pulse waveform anddurations predicted by the use of the present model of effective currentI_(pe) are equally applicable to both cardioversion and defibrillationcountershocks. Typically, cardioversion countershocks are countershocksthat have total pulse energies of between 0.5 and 5 Joules, whereasdefibrillation countershocks are countershocks that have total pulseenergies greater than about 3 Joules. In each case, the presentinvention utilizes cardioversion/defibrillation countershocks which havepulse durations and waveforms optimized in terms of effective currentI_(pe) to produce a smaller total energy required for an effectivecountershock pulse.

Smaller Effective Capacitance Value

For a given capacitor technology, capacitor volume is proportional tothe stored energy. To maximize the performance of an ICD, for a givendisplacement volume, one must therefore optimize the effective currentI_(pe) for a given maximum E_(c) as defined by Eq. 1. Reworking Eq. 1 interms of V_(d) :

    V.sub.d =((2*E.sub.c)/C).sup.0.5                           (Eq. 29)

Combining Eq. 29 with the effective current I_(pe) formula of Eq. 17yields: ##EQU6##

Substituting Eq. 26 for d and Eq. 27 for the optimum tilt gives:##EQU7##

Note that the term for energy can be separated from the remainder of theexpression. Thus, the optimum capacitance value is independent of thestored energy in the capacitor. The numerical solution is:

    (R*C)/d.sub.c =0.795906≈0.8                        (Eq. 32)

The solution is a reasonable result that implies that an RC timeconstant of the countershock pulse should be close to the "timeconstant" of the myocardial cells (i.e., the chronaxie value timeconstant) for optimum performance of the cardioversion/defibrillationcountershock. This solution is accurate to 1% over a broad range of theexogenous variable R and d_(c). The solution also suggests that there isno first order relationship between energy storage and optimumcapacitance. In other word, to change the energy of an ICD, theeffective current model of the present invention suggests that thevoltage should be adjusted up or down, and that the capacitance shouldnot be moved significantly from the ideal value.

Assuming a chronaxie of 2.7 ms and an inter-electrode resistance of 50Ω, the optimum capacitance value from Eq. 32 is 43 μF. Using thiscapacitance value in Eq. 26 yields a pulse duration of 2.83 ms. Thiscorresponds to a tilt value of approximately 73% (see Eq. 27). It willbe noted that the 2.83 ms optimal duration is very close to the 2.7 msassumed chronaxie time. Thus, the optimal pulse duration for thepractical capacitive discharge pulse is close to that of an idealrectangular pulse as represented by the chronaxie time constant,assuming that the capacitance value is optimized to about 43 μF.

As shown in FIG. 10, a presently approved ICD delivers a maximumdefibrillation countershock monophasic pulse with an effective currentI_(pe) of 6.79 Amps from a 140 μF capacitor charged to 750 Volts. Thisrequires a minimum E_(c) of at least 39.4 Joules and a correspondingrequisite volume for this storage. If the optimum capacitor value of 43μF and a tilt of 73% were used, the same 6.79 Amps effective currentI_(pe) could be delivered from a charge of 1195 Volts that would involvean energy storage of only 30.7 Joules. In other words, the lessefficient design of current ICDs with overly large capacitors requires28% more energy and requires 28% more capacitor volume for the device todeliver a monophasic countershock. As previously described, the batteryvolume in existing devices is necessarily larger to supply thisincreased minimum E_(c).

As indicated in the background art section, when designed with presentday microelectronic switches, such as power FETs, the optimal dischargevoltage V_(d) of almost 1200 Volts is problematical. Thus, in thepreferred embodiment a compromise in capacitor size is necessary. As aresult, the optimum capacitance value selected for a discharge voltageV_(d) of 750 Volts, is approximately 85 μF. As a result, the effectivecurrent I_(pe) delivered by the preferred embodiment is 5.67 Amps. Whenthe 13-29% increased efficiency of the effective current of a biphasicwaveform is factored in, the present invention provides a monophasicequivalent effective current of between 6.55 to 7.32 Amps. Thus, thepreferred embodiment that stores a maximum E_(c) of about 26 Joules canactually be more effective than the prior art monophasic defibrillationcountershocks that required a maximum E_(c) of 39.4 Joules, yet onlyproduces an effective current I_(pe) of 6.79 Amps. In addition, thiscompromise on capacitor size of the preferred embodiment permits asimple implementation of the capacitor-discharge ICD, while obtainingthe majority of the benefits of the reduced size of the capacitor, and,the corollary reduction in the size of the battery.

Improved Battery Configuration

For the battery, the physical size of the battery will be primarily afunction of the amp-hours of storage capacity provided by the battery.In addition to the required maximum E_(c), two other energy parametersare required in order to budget the storage capacity of a battery for anICD. These parameters are the minimum number of countershocks that areto be delivered over the life of the device (N_(p)) and the idle currentdrain of the device when it is sensing the cardiac signals (I_(i)).Current ICDs budget for at least 200 defibrillation countershocks overthe life of the device and idle currents are on the order of 15-20μAmps.

Some ICDs also provide for pacing capabilities, in which case therequired pacing energy must also be factored into the storage capacityof the battery. The current draw on a battery due to constant pacing canbe estimated by assuming that the pacing countershock will have a 6 Voltamplitude, a 500 μsec width, and a 500 Ω assumed impedance, and thatpacing will occur at a rate of 70 beats/minute. Under these conditions,the energy drawn from the battery will be about 2.5 mJoules/ minute, oran average current draw of about 7 μAmps. It should be noted that theelectrical energy necessary to provide for pacing capabilities is muchless than even the idle current drawn by the ICD.

If an ICD is designed against an optimum battery budget that wouldsupport a device life (l) of five years, the total storage capacity(E_(t)) required for the battery is the sum of the maximum electricalcharge energies, the maximum idle current energies and the maximumpacing energies. For existing ICDs, such a battery budget can becalculated as follows: ##EQU8##

Most ICDs use a pair of 3 Volt, 2 Amp-hour lithium/silver vanadium oxidebatteries to provide this amount of total storage capacity for the ICD.The lithium/silver vanadium oxide batteries represent the densest powersource technology currently viable for use in an ICD. While improvementsin battery technology may increase the storage density slightly, andthereby decrease the total volume of the power source somewhat, thetotal volume required by the power source for current ICDs must besufficiently large to supply a total storage capacity (E_(t)) for thedevice of about 2.0 Amp-hours.

In contrast to the prior art, the present invention budgets a totalstorage capacity E_(t) for the device of about 1.0 Amp-hours. Thissignificantly smaller storage capacity E_(t) is achieved primarily dueto the smaller maximum E_(c) for the device. Additionally, severalinnovations in capacitor charging and battery configuration of animproved battery system are utilized in the preferred embodiment tofurther reduce both the storage capacity E_(t) and the overalldisplacement volume of the battery system. These innovations include:(1) the use of a dual battery configuration, one battery for themonitoring requirements and a different battery for charging thecapacitor; (2) the use of a staged energy circuit with a rechargeablebattery to charge the capacitor; and (3) the use of an intensifyingbattery to improve the energy delivery capability and decrease thecurrent delivery requirements of the battery system when deliveringmultiple closely spaced defibrillation countershocks, such as during afibrillation incident where the first cardioversion/defibrillationcountershock is not successful.

Dual Battery Configuration

Current ICDs utilize a single battery system to provide all of theenergy storage requirements for the device. Unfortunately, the idealvoltage requirements for the monitoring and capacitor-dischargefunctions of an ICD are almost opposites. For the monitoring function,it is desirable to use the lowest possible voltage that the circuits canoperate reliably with in order to conserve energy. This is typically onthe order of 1.5 to 3.0 Volts. On the other hand, the output circuitworks most efficiently with the highest possible voltages, including upto 800 Volts. The single battery system of current ICDs is typicallycomprised of two lithium vanadium pentoxide cells in series that produceabout have a battery output voltage (V_(b)) of about 6 Volts. Thisvoltage V_(b) is not ideal for either the monitoring orcapacitor-discharge functions.

In the preferred embodiment of the present invention, two separatebattery systems are used to provide the energy storage requirements ofthe ICD, one having optimized characteristics for the monitoringfunctions and one having optimized characteristics for thecapacitor-discharge functions. The preferred battery system is aconventional pacemaker power source for the monitoring functions, suchas a lithium iodide battery, that is optimized for long life at lowcurrent levels. The preferred battery system for the capacitor-dischargefunction is a conventional ICD battery, such as a lithium vandiumpentoxide battery, that is optimized for high current drain capabilityand low self-discharge for long storage life with few discharges. Due toan excess of low current level in the conventional ICD battery used forthe capacitor-discharge function, this battery can also power any pacingfunctions of the device without affecting its operational requirementsto perform the capacitor charging function. By optimizing the twoseparate battery systems, the overall charge density of each battery canbe increased and, hence, the combined volume of both battery systems canbe decreased when compared to the single battery system found in theprior art.

FIG. 16 illustrates a block diagram of the preferred embodiment of thedual battery system 130. A battery 132 of appropriate voltage andminimum physical size connects to and powers a monitoring circuit 134only. Another battery 136 of appropriate voltage and minimum physicalsize connects to and powers the capacitor-discharge output circuit 138only. The monitoring circuit 134 and the capacitor-discharge outputcircuit 138 each connect to electrodes 140 positioned near or in theheart 142. The monitoring circuit 134 also connects to and triggers thecapacitor-discharge output circuit 138 in the event an arrhythmia isdetected. The battery systems 132 and 136 are optimally sizeelectrically and physically to provide for the most efficient operationin the smallest displacement volume.

FIG. 17 illustrates a dual battery system 150 for an ICD where thebatteries are rechargeable. A battery 152 of appropriate voltage andminimum physical size connects to and powers a monitoring circuit 154only. Another battery 160, which is rechargeable and of appropriatevoltage and minimum physical size connects to and powers thecapacitor-discharge output circuit 162 only. Charging of the battery 160occurs by a radio frequency link between an external charger circuit 168and an implanted recharge circuit 170. A coil 172 connects with theexternal charger circuit 168 and transmits RF energy from the coil 172through the epidermis 176 where it is received by the implanted coil174. The coil 174 supplies RF energy to the recharge circuit 170 so thatthe battery 160 may be charged. The dual battery system 150 operates andis sized in a manner similar to the dual battery system 130. In the dualbattery system 150, the ICD has a finite and predictable monitoring lifebased upon the capacity of the primary pacing battery 152, and aninfinite life for the output power surface battery 160 based on atheoretically perfect secondary rechargeable battery. Optionally, thebattery 152 which powers the monitoring circuit 154 could also berechargeable and would include another similar RF charging link as usedfor rechargeable battery 160.

Staged Energy Circuit with an Internally Rechargeable Battery

FIG. 18 is a simplified circuit diagram of a known implantabledefibrillator circuit 210. Circuit 210 comprises a high currentdefibrillation battery 213, which is typically a lithium silver vanadiumpentoxide (LiAgVO₅) battery. A high voltage transformer 215 comprises atransistor switch 218 which drives the primary 221. The oscillatordriving switch 218 provides an alternating current through the primaryof transformer 215. The secondary 225 of transformer 215 produces asignificantly higher voltage which is rectified by diode 227 and storedin capacitor 230. When capacitor 230 is fully charged, the semiconductorswitch 232 is activated to complete the circuit which delivers thecharge of capacitor 230 to the cardiac electrodes 235 for defibrillationof the heart. A configuration which is similar to the above circuitcomprises substitution of a H-bridge in place of switch 232. Thispermits delivery of the current from capacitor 230 in either polarity,which allows delivery of a biphasic pulse.

FIG. 19 discloses a simplified schematic staged energy circuit 240.Circuit 240 comprises a first embodiment of an improved staged energyconcentration means. Circuit 240 preferably comprises a first stage ofenergy concentration comprising a non-rechargeable battery, such as ahigh energy density pacing battery 245, configured for applying a smallμAmp current to the trickle charge control circuitry 248. This providesan optimum current to be supplied to a second stage of energyconcentration, comprising at least rechargeable battery means. Therechargeable battery means preferably comprises a rechargeabledefibrillator battery 250 and is maintained fully charged by the pacingbattery 245. Rechargeable defibrillator battery 250 is used to driveprimary 221 of the high voltage transformer, or similar power transfermeans, through a switch 218 in a manner similar to conventionalcircuits.

The staged energy concentration configuration of circuit 240 permits useof high density pacing batteries to store energy in combination with avery small rechargeable defibrillator battery to deliver a high currentfor up to about 10 shocks. A typical defibrillator will deliver about200 defibrillator shocks. Assuming each of the shocks is less than 30Joules as provided for by the present invention, and with transformerlosses of 25%, the energy system must store 200×240 J=8000 Joules.However, due to this staged energy arrangement, the rechargeable batteryneed only store enough energy for a typical cardiac defibrillationsession of about 5 shocks in less than about four minutes. The fourminute maximum for each defibrillation session is used because braindamage occurs if defibrillation is not successful in about four minutes.The battery means comprising the second stage of energy concentrationmust therefore only store about 5×240 J=200 Joules. Although this isvery little energy, the second stage battery means must be able todeliver a fairly high current of about 1-2 Amps.

FIG. 20 discloses another embodiment of the staged energy concentrationinvention. Circuit 260 discloses a single cell pacing battery 263 whichis used to power a voltage doubler circuit 267. This doubler circuit267, which comprises numerous embodiments, may be configured to producean output of approximately 6 Volts for charging a rechargeabledefibrillation battery, such as battery 270.

Another embodiment of a staged energy concentration defibrillatorcircuit is shown in FIG. 21, in which circuit 276 comprises first stagebattery 280. Battery 280 is a low voltage, for example a 2.8 Volt, LiIsingle cell battery which charges two second stage batteries 283 and284. Batteries 283, 284 are preferably Lithium Titanium Disulfide(LiTiS₂) batteries. Preferably, battery 284 is charged through diode286, battery 283 is charged through diode 287, and resistor 289 is usedwith a preferred value of 10 K Ohms. Field effect transistor switch 292is off during this time. It is recognized that this schematic circuit isfurther simplified because there is optimal trickle charge currentlimiting between battery 280 and the two diodes, however, that detail isnot considered important to this depiction of the invention.

When fibrillation is detected by related detection circuitry, it is thentime to charge the defibrillation capacitor(s) and switch 292 is turnedon. That places batteries 283 and 284 in series, providing a voltage ofapproximately 5 volts for the transformer primary 221. As above,oscillating switch 218 is used to cause a pulsating current to passthrough primary 221 of the transformer.

Use of a multi-stage energy concentration defibrillator, as disclosed inFIGS. 19-21, provides great savings in both volume and weight of thedefibrillator. For example, since the defibrillator battery chemistryhas about half the density of the pacing battery, it is possible toreduce the total battery weight and volume by greater than about 50%.This provides dramatic improvement in the manufacture, implantation, andoperation of the defibrillator, particularly in view of the restrictedsize of desired pectoral implant sites.

The invention further comprises a multi-stage energy concentrationtechnique for a defibrillator in which the defibrillator capacitor meanscomprises either a third stage or a secondary sub-circuit of the secondstage. In either configuration, it is advantageous to provide arechargeable second stage or intermediate battery means as a fullycharged high current output battery means. This permits rapid chargingof the defibrillator capacitor means. Indeed, in certain configurationsit is now possible to recharge at a rapid 3-5 second rate using thisinvention rather than at a slower rate, which is common in the industry.Therefore, yet another advantage of this invention derives from the useof the second stage energy concentration as a recharge rate accelerator.This also results in a defibrillator with reduced end of life chargedegradation due to the constantly recharged second stage. This featureeffectively provides a battery life extension capability before electivereplacement, assuming certain accepted energy levels.

Intensifying Battery

FIG. 22 illustrates representative circuitry for generating a monophasicdefibrillation pulse in the prior art. Circuit 318 comprises battery 320which is used to provide a current through the primary winding oftransformer 323. The current is cycled on and off at a high rate ofspeed by switching transistor 325. The output from transformer 323 isrectified by diode 328 and is captured in the main storage capacitor332. In order to deliver the pulse to the heart, silicon controlledrectifier 336 is triggered providing a current path from capacitor 323to the electrodes 340, 341 in the heart. At the point of pulsetruncation at the end of time period T₁, silicon controlled rectifier344 is triggered. This quickly discharges capacitor 332 and back biasessilicon controlled rectifier 336 to shut off the flow of current throughelectrodes 340, 341 to the heart.

Biphasic and multiple pulse defibrillation countershock have beenexperimented with for many years. Ventricular defibrillation of dogswith waves comprising two pulses with a pulse length and pulse intervaladjusted so that those cells excitable at any moment are defibrillatedby the first pulse and are refractory to the second pulse was disclosedby Kugelberg as early as October 1965, in the Scandinavian Society ofThoracic Surgery, pages 123-128. Kugelberg considered a variety ofpulses and spacings and found that ventricular defibrillation was indeedquite possible with multiple pulses. U.S. Pat. No. 4,996,984, issued toSweeney, discloses adjusting the timing between multiple bursts ofdefibrillation energy based upon the fibrillation cycle length of themammal. Similarly, Sweeney and Reid disclose that the interactionbetween multiple pulses is non-linearly related to the fibrillationcycle length, and that the spacing between multiple pulses may be afixed percentage of the spacing between fibrillation zero crossings inthe heart. (II-610, Supplement II Circulation, Vol. 84, No. 4, October1991, No. 2425). Johnson et al disclose that successive biphasic shocksdelivered through two different electrodes may be either beneficial ordetrimental depending on the delay between the two shocks. (NASPEAbstracts, April 1991, Part II, no. 391; PACE, Vol. 14, p. 715). Otherexamples of multiple pulse defibrillation systems include U.S. Pat. No.5,107,834 to Ideker et al and U.S. Pat. No. 4,708,145 to Tacker, Jr. etal.

In principal, the above disclosures demonstrate that the energy perpulse in a multiple or closely spaced pulse technique using differentpathways, multiple defibrillators, or other inefficient means of energygeneration and distribution, may be reduced from what is commonly usedin a single monophasic or widely spaced pulse technique. Thus, it can beseen that a multiple pulse waveform might lower the total defibrillationenergy threshold by up to 50 percent. However, such a system requiresmultiple sizeable capacitors. The intermediate power intensifier asdisclosed below teaches means for overcoming the impediments of thetheoretical multiple pulse systems. Also disclosed is a novel means forproviding a rapid pulse power system for use with conventional ICDcircuits to permit optional prompt transition from a widely spaceddefibrillation pulse sequence to a closely spaced defibrillation pulsesequence.

The energy generation problem is appreciated more fully by calculatingthe charging power required of a representative capacitor system in anICD device. Assuming a conventional monophasic defibrillator which isdesigned to deliver a 30 Joule countershock, a 12 second delay forcapacitor charging is considered acceptable after fibrillation isdetected. The charging power is described by simple calculation of 30Joules divided by 12 seconds, which yields about 3 watts. This 3 wattlevel of power is available from high quality defibrillation primarycells, such as lithium silver vanadium pentoxide cells, although othersmay be suitable.

Assuming a use of two closely spaced pulses, defibrillation could occurwith 15 Joules in each pulse. The capacitor could be designed to storeonly 15 Joules and could be made of only half the size of presentcapacitors. However, although the capacitor has 12 seconds to charge inorder to create the first pulse by use of present circuitry, thecapacitor then must be quickly recharged to provide the second pulse.Generally, the amount of time required to quickly recharge is the sametime as that required for optimum spacing between the two pulses, whichis about 0.25 seconds. Therefore, the charging power must be equal to 15Joules divided by 0.25 seconds. This requires a 60 watt power source.Currently, there is no functional implantable battery which is capableof providing such power output.

FIG. 23 discloses the essential circuit elements of one embodiment ofthe present invention in which circuit 367 uses both a primary batteryand an intermediate power intensifying battery, with the lattercomprising a very high power output battery to provide the high chargingpower between capacitor pulses. As shown, battery 370 is a low amperageprimary defibrillation cell, which is preferably a lithium silvervanadium pentoxide type, although other materials are feasible. Whenfibrillation is detected, battery 370 is used to quickly charge arechargeable battery 373 which is capable of very high power output.This is preferably accomplished through the use of transistor switch376. Battery 373 is preferably selected from a list of possible highpower rechargeable batteries, such as a lithium titanium disulfide,lithium sulphur dioxide, or others suitable for producing the desiredpower in a rechargeable configuration. When battery 373 has beensufficiently charged then it is useful as a source of high currentcharging power to capacitor 332 in circuitry sub-section 382, shown incircuit 318 of FIG. 22 and in circuit 367 of FIG. 23.

The difference between the capacitor charging circuitry of FIG. 23 andFIG. 22 is an approximately 20:1 charging power ratio of 60 watts ratherthan 3 watts. The charging circuitry shown as schematic circuit 367provides power means for recharging the capacitor of the related ICDdevice, after an initial discharge, between subsequent multiple pulses.This eliminates additional capacitors and eliminates about half of thecapacitor volume of known ICD devices. The invention also results insignificant improvement in size and operation of an ICD device.

Another alternate embodiment for charging rechargeable battery 73 afterfibrillation is detected is disclosed in FIG. 24. In this embodiment,circuit 390 comprises a relatively low amperage, e.g. milliamps, primarydefibrillation battery 393. One example of such a battery 393 comprisesa pacing type lithium iodide battery, although other materials are alsosuitable. Circuit 390 also comprises high power output (approximately1-3 Amps) intermediate power intensifying battery 397. A preferredbattery 397 comprises a lithium titanium di-sulfide battery. Battery 301comprises a very high amperage (10-30 amps) battery. In operation,circuit 390 allows continuous trickle charge from battery 393 to battery397. This maintains battery 97 in a substantially fully chargedconfiguration until detection of fibrillation. After detection offibrillation, battery 397 simultaneously charges the main energydelivery capacitor 332 within sub-section 382 and battery 301, viaswitch 305. Capacitor 332 then discharges and is again re-charged withbattery 397. However, battery 397 is not normally able to fully chargecapacitor 332 in less than at least about 5 seconds. In a closely spacedmultiple pulse ICD device power system it is necessary to provide meansother than battery 397 to provide charging power for subsequent pulsesto the heart. Rather than providing multiple charging pathways or aplurality of capacitors, circuit 90 discloses use of battery 301 toprovide high amperage high power means for charging a main energydelivery capacitor for countershock pulses after the initialcountershock/pulse.

The embodiment of circuits 367 and 390 are each also advantageous as arapid pulse power system for use with implantable cardioverter devices.This rapid pulse power system may be integrated into other known orproprietary circuits as a means of enabling rapid and optionaltransition from a widely spaced defibrillation pulse sequence to aclosely spaced defibrillation pulse sequence. This is accomplishedwithout adding any additional capacitors, which would detract from thesize and volume advantages of the invention.

Optimized Battery Budget

Budgeting the E_(t) for the preferred embodiment of the ICD of thepresent invention, it will first be noted that the lower maximum E_(c)of the present invention produces a minimum charging time (t_(min)) of8.5 seconds, in contrast to the 12 second minimum in the prior artdevices. The use of the improvements to the battery system as previouslyset forth allows the idle current I_(i) and the pacing current I_(p) tobe reduced by about half to about 10 μAmps and 3.5 μAmps, respectively,as compared to about 20 μAmps and 7 μAmps in the prior art devices. Byreducing the budgeted number of countershock N_(p) to about 150, ratherthan 200, it will be seen that the E_(t) of the preferred embodiment iseffectively reduced in half as compared to the prior art devices.##EQU9##

Although the description of the preferred embodiment has been presented,it is contemplated that various changes could be made without deviatingfrom the spirit of the present invention. Accordingly, it is intendedthat the scope of the present invention be dictated by the appendedclaims, rather than by the description of the preferred embodiment.

What is claimed is:
 1. An implantable cardioverter defibrillator forsubcutaneous positioning within a pectoral region of a human patientcomprising:a sealed housing structure constructed of a biocompatiblematerial; a battery source of electrical energy contained in thehousing; a capacitor system connected to the battery source wherein thecapacitor system stores electrical energy provided by the battery sourceto generate high voltage electrical cardioversion/defibrillationcountershocks; and a control system connected to the capacitor system tocontrol the storing of the electrical energy in the capacitor system andto control the discharge of the electric energy in the capacitorsystems; and wherein a total mass of the implantable cardioverterdefibrillator is less than about 120 grams.
 2. The implantablecardioverter defibrillator of claim 1 wherein a maximum electricalcharge energy stored by the capacitor system for each electricalcardioversion/defibrillation countershock is less than about 30 joules.3. The implantable cardioverter defibrillator of claim 1 wherein thebattery source charges the capacitor system to the maximum electricalcharge energy in less than about 10 seconds.
 4. The implantablecardioverter defibrillator of claim 1 wherein the capacitor system hasan effective capacitance value of less than 120 μF.
 5. The implantablecardioverter defibrillator of claim 1 wherein a maximum electricalcharge energy stored by the capacitor system for each electricalcardioversion/defibrillation countershock is less than about 30 joules,wherein the battery source charges the capacitor system to the maximumelectrical charge energy in less than about 10 seconds, and wherein thecapacitor system has an effective capacitance value of less than about120 μF.
 6. The implantable cardioverter defibrillator of claim 5 whereina total volume of the sealed housing is less than about 60 cubiccentimeters.
 7. The implantable cardioverter defibrillator of claim 1wherein a total volume of the sealed housing is less than about 60 cubiccentimeters.